This invention relates to nuclear magnetic resonance (NMR) imaging methods and systems. More specifically, this invention relates to the tuning of radio frequency (RF) coils used with NMR apparatus to transmit and receive RF signals.
In NMR imaging, a uniform magnetic field B.sub.0 is applied to the imaged object along the z axis of a Cartesian coordinate system, the origin of which is centered within the imaged object. The effect of the magnetic field B.sub.0 is to align the object's nuclear spins along the z axis. In response to an RF magnetic signal of the proper frequency, oriented within the x-y plane, the nuclei precess about the z-axis at their Larmor frequencies according to the following equation: EQU .omega.=.gamma.B.sub.0
where .omega. is the Larmor frequency, and .gamma. is the gyromagnetic ratio which is constant and a property of the particular nuclei. Water, because of its relative abundance in biological tissue and the properties of its nuclei, is of principle concern in such imaging. The value of the gyromagnetic ratio .gamma. for water is 4.26 khz/gauss and therefore in a 1.5 Tesla polarizing magnetic field B.sub.0 the resonant or Larmor frequency of water is approximately 63.9 Mhz.
In a typical imaging sequence, the RF signal centered at the Larmor frequency .omega., is applied to the imaged object by means of a radio frequency (RF) coil. A magnetic field gradient G.sub.z is applied at the time of this RF signal so that only the nuclei in a slice through the object along the x-y plane, which have a resonant frequency .omega., are excited into resonance.
After the excitation of the nuclei in this slice, magnetic field gradients are applied along the x and y axes. The gradient along the x axis, G.sub.x, causes the nuclei to precess at different resonant frequencies depending on their position along the x axis, that is, G.sub.x spatially encodes the precessing nuclei by frequency. Similarly, the y axis gradient, G.sub.Y, is incremented through a series of values and encodes y position into the rate of change of phase as a function of gradient amplitude, a process typically referred to as phase encoding.
A weak RF signal generated by the precessing nuclei may be sensed by the RF coil and recorded as an NMR signal. From this NMR signal, a slice image may be derived according to well known reconstruction techniques. An overview NMR image reconstruction is contained in the book "Magnetic Resonance Imaging, Principles and Applications" by D. N. Kean and M. A. Smith.
Referring to FIG. 1, a nucleus 10 has a magnetic moment 12 which may be excited into precession 18 about a static magnetic field B.sub.0 along axis 16 by an RF magnetic signal producing magnetic field 14 along a plane perpendicular to the static magnetic field B.sub.0.
The exciting RF magnetic field 14 may oscillate along a single axis within the x-y plane. Such an oscillating field may be generated by a "saddle" coil (not shown) comprised of two conductive loops disposed along the axis of oscillation and perpendicular to the static magnetic field B.sub.0 as is known in the art.
A more effective excitation of the nuclear moments 12 may be achieved with a circularly polarized RF magnetic field 14, i.e. one that produces a rotating magnetic field vector 14. Preferably, the magnetic vector of field 14 rotates within the x-y plane at an angular velocity equal to the Larmor frequency .omega. as shown by arrow 20 in FIG. 1.
It is known that a rotating RF magnetic field 14 may be generated with certain RF coil structures when the coil structure is excited at its "resonant" frequency. Referring to FIG. 2(a), a coil structure 28 for creating a rotating magnetic field is comprised of a pair of conductive loops 22 and 22' spaced along the axis 16 of the static magnetic field B.sub.0. The loops 22 and 22+ are joined with a series of conductive segments 24 parallel to axis 16 of the static magnetic field B.sub.0. The loops 22 and 22' and conductive segments 24 have an intrinsic inductance and may be broken along their length with capacitive elements C.sub.e to promote the desired pattern of current flow through the conductive segments 24 when the coil is driven by an external RF generator (not shown). The capacitive elements C.sub.s may be positioned along the conductive segments 24 between the loops 22 and 22' as in FIG. 2(a) in a "low-pass" configuration; along the loops 22 and 22' between the conductive segments 24 as in FIG. 2( b) in a "high-pass" configuration; or in both of the aforementioned positions as in FIG. 2(c) in a "bandpass" configuration.
Detailed descriptions of the above RF coil structures 28 are given in the following U.S. Pat. Nos. assigned to the assignee of the present application and hereby incorporated by reference: 4,680,548, entitled: "Radio Frequency Field Coil for NMR" and issued July, 14, 1987; 4,692,705, entitled: "Radio Frequency Field Coil for NMR" and issued Sept. 8, 1987; and, 4,694,255, entitled: "Radio Frequency Field Coil for NMR" and issued Sept. 15, 1987. These designs will be referred to collectively as "resonant RF coils".
When a resonant RF coil 28 is driven in a first resonant mode at a particular frequency f.sub.1, the phase of the current distribution in each axial segment 24 will be proportional to the transverse angle .theta. of the segment 24 measured around the axis 16 of the static magnetic field B.sub.0. This phase distribution is the result of a "delay line" effect of the distributed inductance of the coil 28 in combination with the capacitive elements C.sub.e and C.sub.s which together create a standing wave of voltage along each loop 22 and 22'. At the driving frequency f.sub.1 the delay line produces a full .theta. of phase shift in the current flowing though the conductive segments 24, for 360.degree. of angular displacement .theta. of the conductive segments 24. As is understood in the art, this current distribution produces an RF a magnetic field 14 having a sinusoidally varying magnitude directed along a first fixed axis at some angle .theta..sub.1.
By appropriately driving the coil structure 28 at a second mode with a second signal of frequency f.sub.2 =f.sub.1, a second orthogonal standing wave, angularly displaced along the loop 22 and 22' by 90.degree., may be generated. This second standing wave will create a second magnetic field 14 having a sinusoidally varying magnitude directed along a second fixed axis at some angle .theta..sub.2 orthogonal to .theta..sub.1. The combined effect of the two modes will be to create a circularly polarized RF magnetic field 14 rotating about the longitudinal axis 16 with an an angular velocity equal to the frequency of the first and second driving signals f.sub.1 and f.sub.2.
Typically, these two orthogonal modes are generated by exciting the coil structure 28 at two different excitation points displaced from one another by 90.degree. about the longitudinal axis 16. The coil structure 28 may be driven by a RF generator (not shown) directly connected across one of the capacitive elements C.sub.s in an conductive segment 24. Alternatively, U.S. Pat. No. 4,638,253, entitled: "Mutual Inductance NMR RF Coil Matching Device," issued Jan. 20, 1987, teaches a method of inductively coupling an RF source 26 to the coil structure 28. This patent is also assigned to the assignee of the present application and hereby incorporated by reference.
If the values of the capacitances and inductances distributed around the coil structure 28 are not equal, the two orthogonal resonant modes of the coil structure may have different frequencies f.sub.1 and f.sub.2 oriented at angles .theta..sub.1 and .theta..sub.2. As a rule, the fields 14 generated by each mode will remain orthogonal regardless of their absolute orientation and their frequencies, and will shift to an angle about the longitudinal axis 16 that permits the greatest difference between f.sub.1 and f.sub.2 as a result of variations in capacitance and inductance around the coil structure 28. One result of this is that only .theta..sub.1 need be measured as .theta..sub.2 will always equal .theta..sub.1 +90.degree..
If the two resonant modes of the coil structure have different frequencies f.sub.1 and f.sub.2, the uniformity of the RF magnetic field 14 will be degraded and the efficiency of energy transfer between the coil and the NMR apparatus will be decreased.
Small variations in the values of the inductances and capacitances of the coil structure 28 may be "tuned" out by the use of reactive shunts positioned around the circumference of the coil. The shunts may be used to trim the reactive elements of the RF coil and thus adjust f.sub.1 =f.sub.2.
Determining the proper value of each shunt necessary to tune the RF coil 28 is accomplished by measuring f.sub.1 and f.sub.2 of the untuned RF coil 28 by exciting the coil 28 through a range of frequencies with a signal generator and using an inductively coupled network analyzer to plot the strength of the current flow in the coil 28 versus the excitation frequency at various locations about the coil 28. The plot will show two closely spaced peaks corresponding to the resonant modes of frequency f.sub.1 and f.sub.2. The location on the RF coil 28 with the greatest current flow for one peak, as determined by moving the inductive coupling of the network analyzer to various places on the coil 28, will indicate the location of one of either .theta..sub.1 or .theta..sub.2 from which the other may be determined.
This tuning method may be difficult and time consuming. Often the peaks at f.sub.1 and f.sub.2 are not clearly differentiated or obscured by noise. The requirement that the inductive coupling for the signal generator and network analyzer be moved around the coil 28 is complicated by the fact that network analyzer must be isolated by a significant distance from the large magnetic field associated with the NMR system and hence each move requires the operator to walk back and forth a distance between the equipment and the coil.